Embodiments of the invention relate generally to a scintillator for x-ray detection, and more particularly to a rare earth garnet scintillator composition having enhanced performance and a method for making same.
Typically, an imaging system such as a computed tomography (CT) or an x-ray imaging system includes an x-ray source positioned to emit x-rays toward a detector, and an object positioned therebetween. In CT imaging systems, the x-ray source emits a fan-shaped beam toward a subject or object, such as a patient or a piece of luggage. Hereinafter, the terms “subject” and “object” shall include anything capable of being imaged. Generally, the x-ray source and the detector array are rotated about the gantry within an imaging plane and around the subject. X-ray sources typically include x-ray tubes, which emit the x-ray beam at a focal point.
The x-ray beam, after being attenuated by the subject, impinges upon an array of x-ray detectors. The intensity of the attenuated beam radiation received at the detector array is typically dependent upon the attenuation of the x-ray beam by the subject. X-ray detectors typically include a collimator for collimating x-ray beams received at the detector, a scintillator for converting x-rays to light energy adjacent the collimator, and photodiodes for receiving the light energy from the adjacent scintillator and producing electrical signals therefrom. In order to achieve high-resolution images, the scintillator material is often diced into small pieces or cells and assembled in a pixilated array with desired geometry prior to attaching to the photodiode.
The scintillator material of a cell or element of a scintillator array converts x-rays to light energy. In particular, the scintillator material absorbs x-rays incident on that cell and discharges light energy (photons) to a photodiode adjacent thereto. Each photodiode detects the light energy and generates a corresponding electrical signal indicative of the attenuated beam received by each detector element. The outputs of the photodiodes are then transmitted to the data processing system for image reconstruction.
Scintillators typically comprise a non-luminescent host material that has been modified by inclusion of an activator species that is present in the host material in a relatively low concentration. The host crystal absorbs the incident photon, and the absorbed energy may be accommodated by the activator ions, or it may be transferred by the lattice to the activator ions. One or more electrons of the activator ions are raised to a more excited state. These electrons, in returning to their less excited state, emit a photon of luminescent light.
Solid state ceramic scintillators are currently used as radiation detectors to detect penetrating radiation. One embodiment of the present generation of solid-state ceramic scintillators comprises oxide mixtures in which a rare earth oxide is present as an activator, along with various combined matrix elements which are also usually rare earth oxides. One embodiment of the present generation of solid-state ceramic scintillators comprise a major proportion of yttria (Y2O3), up to about 50 mole percent gadolinia (Gd2O3), and a minor activating proportion (typically about 0.02-12 mole percent) of a rare earth activator oxide. Activator oxides include europium, neodymium, ytterbium, dysprosium, terbium, cerium, and praseodymium. Other metallic compounds may also be present as additives for specific purposes.
The material properties of scintillators vary greatly based on the specific chemical composition of the scintillator. These properties include scintillator efficiency, primary decay time, afterglow, hysteresis, luminescent spectrum, x-ray stopping power, and resistance to radiation damage. The efficiency of a luminescent material is the percentage of the energy of the absorbed stimulating radiation which is emitted as luminescent light. When the stimulating radiation is terminated, the luminescent output from a scintillator decreases in two stages. The first of these stages is a rapid decay from the full luminescent output to a low, but normally non-zero, value at which the slope of the decay changes to a substantially slower decay rate. This low intensity, normally long decay time luminescence, is known as afterglow. Specifically, afterglow is the light intensity emitted by the scintillator at 100 milliseconds after the x-ray excitation ceases, reported as a percentage of the light emitted while the scintillator is excited by the radiation. Afterglow provides a background luminescent intensity, which is a noise contribution to the photodetector output. In some cases, afterglow is increased by the presence of impurities, and in other cases, afterglow is decreased by the presence of impurities.
The initial, rapid decay is known as the primary decay or primary speed and is measured from the time at which the simulating radiation ceases to the time at which the luminescent output falls to about 36.8% (or 1/e) of the light intensity at the time after the x-ray excitation ceases. Scan times of CT systems, which are the times required for a CT system to scan and acquire an image of a slice of the subject under observation, are related to the primary decay time of the scintillator roughly by a factor of 1,000. For example, a scintillator having a decay time of 1 millisecond will typically produce a scan time of about 1 second. Thus, shorter CT scan times require shorter scintillator decay times. As the speed of data processing in CT scanners increases due to advances in electronic circuit designs, it is desired to have faster scintillators, i.e., shorter time between receipts of stimulating radiation pulses so to fully take advantage of the capability of the scanner. Therefore, any measurable percentage decrease in decay time from that exhibited by the present generation of ceramic scintillator would be a distinct improvement. Decreasing scan time increases the number of patients that can be scanned, as well as the number of scans taken in a single measurement, as each measurement requires a patent breathold during the measurement period. Shorter scan times also reduce image blurring due to motion of internal organs or motion that occurs when taking scans of non-cooperating patients, such as young children.
Another important consideration for scintillators is to reduce damage that occurs on the scintillator upon repeated exposure of the scintillator to high energy radiation. Radiographic equipment employing solid state scintillator materials for the conversion of high energy radiation into an optical image may experience changes in efficiency after exposure of the scintillator to high dosages of radiation. Radiation damage is the characteristic of a luminescent material in which the quantity of light emitted by the luminescent material in response to a given intensity of stimulating radiation changes after the material has been exposed to a high radiation dose. Radiation damage in scintillators is characterized by a change in light output and/or a darkening in color of the scintillator body with prolonged exposure to radiation. Radiation damage can lead to “ghost images” from prior scans which thereby reduce image resolution. When radiation damage becomes too high, the scintillator must be replaced because of the cumulative effects of the radiation damage. This results in a substantial capital cost for the replacement of the scintillation detecting system. Additionally, the effects of radiation damage may require recalibration of the imaging system throughout the work day. Such recalibration takes time and also exposes the scintillator material to additional radiation, which contributes further damage.
In systems such as CT scanners, it is also desirable that the scintillator have high x-ray stopping power, which refers to the ability of a material to absorb radiation, commonly called the attenuation or absorption. A material having a high stopping power allows little or no radiation to pass through. Since x-ray stopping power is inversely related to the patient radiation dose required for obtaining high quality patient images, a scintillator with high x-ray stopping power can absorb substantially all of the incident x-rays in the luminescent material in order to minimize the x-ray dose to which the patient must be exposed in order to obtain the CT image. Higher stopping power is also preferred because a smaller quantity of scintillator material is needed. Thus, thinner detectors are possible, resulting in lower cost of manufacture.
In typical medical and industrial radiographic applications, the scintillator also must be an efficient converter of x-ray radiation (or other high-energy radiation) into optical radiation for the regions of the electromagnetic spectrum detected by the photodiode. As used herein, the term “light output” is the quantity of visible light emitted by a respective scintillator element after being excited by a pulse of x-ray.
Thus, there is a need to develop scintillator materials that have a short decay time (to minimize scan time), reduced afterglow, high light output, show reduced damage upon repeated exposure to the high energy radiation typically employed in applications requiring scintillators (to increase the reproducibility of the measurements), and high x-ray stopping power. There is also a need to develop scintillators that are cost-effective to manufacture.
As stated above, the properties of the scintillator vary greatly based on the material composition of the scintillator. Thus, the substitution of one rare earth element for another or even a minute change in the concentration of one element in the scintillator's composition can have significant consequences on the performance of the scintillator. Although work has been done to identify how certain combinations of elements affect scintillator performance, exactly how or why specific elements interact to influence scintillator performance is not fully understood. Thus, the current state of the art is unable to precisely predict the resulting properties of a scintillator from the chemical composition alone.
Therefore, it would be desirable to develop a scintillator that has decreased afterglow, improved efficiency, enhanced stability under radiation, and desirable x-ray stopping power.